In PET imaging a radiotracer is administered to a subject such as a patient or an animal prior to its positioning in the imaging region of a PET imaging system. The radiotracer is absorbed by regions in the subject and its distribution is imaged following an uptake period. Subsequently a clinician interprets images showing the relative uptake of the radiotracer at particular sites and may perform a diagnosis of the subject. The radiotracer undergoes radioactive decay which results in the production of positrons. Each decay event produces one positron which travels up to a few mm in human tissue, where it subsequently interacts with an electron in an annihilation event that produces two oppositely-directed gamma photons. The two gamma photons each have an energy of 511 keV and are detected by gamma photon detectors disposed radially around the imaging region which each produce an electrical signal when struck by an incident gamma photon. The resulting electrical signals are processed by coincidence-detection circuitry which, through the position of the detectors, determine a line in space along which the annihilation event occurred. Gamma photons received within +/−3 ns of each other are typically accepted as being coincident. The endpoints of this line are defined by the two positions at which the coincident events are detected and the line is termed a Line Of Response (LOR). Such LORs are subsequently reconstructed to produce a two- or three-dimensional image illustrative of the distribution of radiotracer within the imaging region.
In time-of-flight (TOF) PET the small time difference between the two detected events is further used to localise the position along the LOR at which the annihilation event occurred, and thus more accurately locate the spatial position of the radiotracer causing the decay event. In depth-of-interaction (DOI) PET, multi-layered detectors located at different radial distances from the imaging region further detect the depth at which the gamma photon is absorbed. In DOI PET this information is used to further improve the spatial resolution of detection by reducing parallax errors.
In PET imaging systems a gamma photon detector is defined hereinafter to comprise a scintillator material and an optical detector. The scintillator material creates a pulse of scintillation light when struck by a gamma photon, and the optical detector, which is optically coupled to the scintillator material, converts the pulse of scintillation light into an electrical signal. When a gamma photon strikes the scintillator material, probabilistic events determine the depth at which the scintillation light is generated, at which point it transfers its energy to the scintillator and the pulse of scintillation light having both a characteristic wavelength spectrum and a characteristic decay time is created. The scintillator material is further characterised by an absorption depth within which a proportion 1/e of the received gamma photons are absorbed. Due to the high energy of the incident gamma photons, dense scintillator materials are preferred in order to absorb a high proportion of incident gamma photons within a practical depth of scintillator material.
Owing to the process of determining the spatial position of radioactive decay events through coincidence, a gamma photon detector in a PET imaging system must be capable of discriminating between the incidence of individual gamma photons. A key parameter which characterises this ability is the maximum gamma photon detection rate. The ability to measure high incidence, or count rates, is desirable in the measurement of images with high signal to noise ratio within short acquisition times. Short acquisition times are important in the prevention of patient motion-induced artifacts in the images. The maximum gamma photon detection rate is affected by the decay of the scintillation light. The decay time of the scintillation material determines the minimum time interval between consecutively-incident gamma photons after which their scintillation light no longer overlaps. Such overlapping events, termed pile-up, must be prevented because they inhibit the ability to count the reception of individual photons. The need to reduce the decay time in PET scintillator materials is further driven by the demand for good timing resolution in TOF-PET. The state of the art in decay time is currently 25 ns in LaBr3 with current research efforts emphasizing the need to reduce this even further.
Light yield and energy resolution are two further scintillator material parameters that characterise a gamma photon detector, particularly in a PET imaging system. The light yield from a scintillator material is the number of scintillation photons that are produced by an incident gamma photon. Light yield is typically normalized to the energy of the gamma photon and expressed as the number of photons produced per MeV. A high light yield, thus a sensitive scintillation material is desirable in the provision of a high signal to noise ratio gamma photon detector since it provides the associated optical detector with a strong light pulse in response to each incident gamma photon. Scintillator materials with good energy resolution provide an additional means of verifying that two photons detected within a narrow time interval indicate a valid LOR. By rejecting events that lie outside a predefined energy window, a PET imaging system may discriminate between scattered gamma photons whose trajectories have been altered by intervening matter and which have energies that lie outside the window, and gamma photons indicative of a valid LOR. One method of providing such discrimination is to determine the energy of each received gamma photon by integrating the scintillation light pulse, and to only accept it as being a valid coincidence event if it is both detected within a narrow time interval of another gamma photon, as well as if its energy is within a narrow energy window of non-scattered gamma photons. Good energy resolution is provided through the use of materials having a large value of effective atomic number. When determining the energy of a gamma photon in this way, pile-up must again be prevented by ensuring the scintillation light decays to a level where it no longer interferes with that from a subsequent gamma photon, and this again requires a short decay time.
In summary, the design of a gamma photon detector, particularly that used in a PET imaging system is driven fundamentally by the need to acquire high quality images with which a clinician can make an accurate diagnosis of a subject. High quality images, or more specifically high signal to noise ratio images demand a sensitive gamma photon detector which also meets the fast timing constraints associated with minimising the duration of the imaging process. This places a number of constraints on the gamma photon detector's scintillator material and optical detector. These are driven primarily by the need for a scintillator material with a short decay time. Providing the light yield from the scintillator material is sufficiently high to give an acceptable signal to noise ratio, an optical detector is subsequently optimised such that its responsivity is improved within the region of optical wavelengths emitted by the scintillator material. Typically, photomultiplier tube (PMT) detectors are used as the optical detector in what is termed analogue PET, and more recently solid state semiconductor optical detectors, defined herein as optical detectors produced using monolithic processes in semiconductors have been used to provide a more integrated system, in what is termed digital PET.
The shortest scintillator decay times in scintillator materials suitable for use in gamma photon detectors are conventionally found in blue-emitting scintillator materials (see for instance: Luminescence: From Theory to Applications, Wiley-VCH, Darmstadt, 2007, C. Ronda (Ed.)). Consequently blue-emitting scintillator materials are preferred, and the associated optical detector, typically a PMT in analogue PET imaging systems, is optimised to provide high sensitivity at around 420 nm wavelength, the emission peak of NAI:TI. Bi-alkali photocathode tubes are typically used since they are commercially available. Blue-sensitive photomultiplier tubes are preferred furthermore because although green/yellow sensitive photomultipliers having multi-alkali photocathodes are available, their lower quantum efficiency renders them less efficient.
Known scintillator materials for PET (Luminescence: From Theory to Applications, Wiley-VCH, Darmstadt, 2007, C. Ronda (Ed.)) include LYSO, LaBr3 and the broad group of materials known as garnets (US2006/0219927A1). Scintillation in LYSO has been reported with a light yield of 33000 photons/MeV in the presence of a decay time of 44 ns, a high density of 7.1 g/cm3, and an energy resolution of approximately 10%, defined as the ratio of the FWHM of the energy detection peak to the peak detection energy. In LaBr3 a decay time of 25 ns has been reported with an improved energy resolution of 3% and higher light yield than LYSO. In the single crystal garnet Ce:Gd3Al2Ga3O12 a light yield of 35000 photons/MeV with a 68 ns decay time has been reported for 1% cerium concentration (K. Kamada et al: 2 inch diameter single crystal growth and scintillation properties of Ce:Gd3Al2Ga3O12. Journal of Crystal Growth 352, 2012, 88-90).
US2012/0223236A1 discloses ceramic garnet compositions according to the composition (Lu, Gd)3(Al, Ga)5O12. In one example composition (Gd0.497Lu0.497Ce0.006)3.04(Al0.6Ga0.04)5O12.06 a decay time of approximately 40 ns is reported.
JP2012-180399 discloses a number of garnet compositions according to the composition Gd3-x-yCexREyAl5-zMzO12 wherein M may be Ga and RE, a rare earth, may be Lu. Crystalline compositions are observed to give rise to a high light yield of up to 68000 photons/MeV. The one disclosed ceramic composition has a light yield of 28000 photons/MeV.
JP2012-066994A discloses a number of single crystal garnet compositions according to the composition Gd3-x-yCexREyAl5-zGazO12 wherein RE may be Lu.
As in the above citations, single crystal materials are investigated almost exclusively owing to the best combination of stopping power, decay time and light yield being found in single crystal materials.